Spinal cord stimulator system

ABSTRACT

A wireless charger for automatically tuning an optimum frequency to inductively charge a rechargeable battery of an implantable pulse generator (IPG) that generates spinal cord stimulation signals for a human body is provided. The charging coil in the charger is wirelessly coupled to a receiving coil of the IPG to charge the rechargeable battery. An optimization circuit detects a reflected impedance of the charging coil through a reflected impedance sensor, and select an optimum frequency of a charging signal supplied to the charging coil based on the detected reflected impedances of a plurality of charging frequencies in a selected frequency range. Advantageously, the optimum charging frequency provides a more efficient way to charge the IPG&#39;s rechargeable battery.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is a continuation of U.S. patent applicationSer. No. 14/642,822, filed on Mar. 10, 2015 (published as U.S. PatentPublication No. 2015-0180271), which is a continuation-in-part of U.S.patent application Ser. No. 14/173,510, filed Feb. 5, 2014 (now U.S.Pat. No. 9,440,076), which claims priority to U.S. ProvisionalApplication Ser. No. 61/792,654, filed Mar. 15, 2013, and titled “SPINALCORD STIMULATOR SYSTEM” (expired), all of which are herein incorporatedby reference in their entirety.

TECHNICAL FIELD

This disclosure relates to stimulators using electrical pulses in amedical context, and more particularly, applying electrical pulsestimulators to the spinal cord to control pain.

BACKGROUND

A Spinal Cord Stimulator (SCS) is used to exert pulsed electricalsignals to the spinal cord to control chronic pain. Spinal cordstimulation, in its simplest form, comprises stimulating electrodesimplanted in the epidural space, an implanted pulse generator (IPG)implanted in the lower abdominal area or gluteal region, conductingwires connecting the electrodes to the electrical pulse generator, anelectrical pulse generator remote control, and an electrical pulsegenerator charger. Spinal cord stimulation has notable analgesicproperties and, at the present, is used mostly in the treatment offailed back surgery syndrome, complex regional pain syndrome andrefractory pain due to ischemia.

Electrotherapy of pain by neurostimulation began shortly after Melzackand Wall proposed the gate control theory in 1965. This theory proposedthat nerves carrying painful peripheral stimuli and nerves carryingtouch and vibratory sensation both terminate in the dorsal horn (thegate) of the spinal cord. It was hypothesized that input to the dorsalhorn of the spinal cord could be manipulated to “close the gate” to thenerves. As an application of the gate control theory, Shealy et al.implanted the first spinal cord stimulator device directly on the dorsalcolumn for the treatment of chronic pain in 1971.

Spinal cord stimulation does not eliminate pain. The electrical impulsesfrom the stimulator override the pain messages so that the patient doesnot feel the pain intensely. In essence, the stimulator masks the pain.A trial implantation is performed before implanting the permanentstimulator. The physician first implants a trial stimulator through theskin (percutaneously) to perform stimulations as a trial run. Because apercutaneous trial stimulator tends to move from its original location,it is considered temporary. If the trial is successful, the physiciancan then implant a permanent stimulator. The permanent stimulator isimplanted under the skin of the abdomen with the leads inserted underthe skin and subcutaneously fed to and inserted into the spinal canal.This placement of the stimulator in the abdomen is a more stable,effective location. The leads, which consist of an array of electrodes,can be percutaneous type or paddle type. Percutaneous electrodes areeasier to insert in comparison with paddle type, which are inserted viaincision over spinal cord and laminectomy.

From time to time, the battery in the IPG needs to be charged wirelesslysince the IPG is implanted in the patient's body. There are a number ofproblems that exist in currently available wireless chargers for theIPG. Problems include inefficient charging, improper charger alignment,difficulty of aligning the charger by patients and lack of ability forthe charger to terminate charging when it is completed. Therefore, itwould be desirable to provide a system and method for an improvedcharger for the SCS system.

SUMMARY

According to one aspect of the present invention, a wireless charger forautomatically tuning an optimum frequency to inductively charge arechargeable battery of an IPG is provided. The charging coil in thecharger is wirelessly coupled to a receiving coil of the IPG to chargethe rechargeable battery. An optimization circuit detects a reflectedimpedance of the charging coil through a reflected impedance sensor, andselect an optimum frequency of a charging signal supplied to thecharging coil based on the detected reflected impedances of a pluralityof charging frequencies in a selected frequency range. Advantageously,the optimum charging frequency provides a more efficient way to chargethe IPG's rechargeable battery.

According to another aspect of the present invention, a method for awireless charger to automatically tune an optimum frequency toinductively charge a rechargeable battery of an IPG is provided. Themethod applies a plurality of charging frequencies in a selectedfrequency range to a charging coil. For each applied charging frequencyof the plurality of charging frequencies, a reflected impedance of thecharging coil is detected. Based on the detected impedances, the methodselects an optimum frequency of the charging coil.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts various components that can be included in a spinal cordstimulation system, according to an embodiment, during trial andpermanent implantation.

FIG. 2 depicts an exploded view of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 3 depicts a feedthrough assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 4 depicts a lead contact system of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 5 depicts a lead contact assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 6 depicts a head unit assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 7 depicts an RF antenna of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 8 depicts a percutaneous lead, according to an embodiment.

FIG. 9 depicts a paddle lead, according to an embodiment.

FIG. 10 depicts a lead extension, according to an embodiment.

FIG. 11 depicts a lead splitter, according to an embodiment.

FIG. 12 depicts a sleeve anchor, according to an embodiment.

FIG. 13 depicts a mechanical locking anchor, according to an embodiment.

FIG. 14 illustrates communication via a wireless dongle with atablet/clinician programmer and smartphone/mobile/patient programmerduring trial and/or permanent implantation, according to an embodiment.

FIG. 15 depicts a Tuohy needle, according to an embodiment.

FIG. 16 depicts a stylet, according to an embodiment.

FIG. 17 depicts a passing elevator, according to an embodiment.

FIG. 18 depicts a tunneling tool, according to an embodiment.

FIG. 19 depicts a torque wrench, according to an embodiment.

FIG. 20 is a function block diagram of a wireless charger according toan embodiment.

FIG. 21 is a function block diagram of an implantable pulse generatoraccording to an embodiment.

FIG. 22 is a functional block diagram of a Class-E amplifier of thewireless charger according to an embodiment.

FIG. 23 is a functional block diagram of a reflected impedance sensor ofthe wireless charger according to an embodiment.

FIG. 24 is a flowchart for a method of optimizing the charging frequencyof the wireless charger according to an embodiment.

DETAILED DESCRIPTION Implantable Pulse Generator (IPG)

FIG. 1 illustrates various components that can be included in a SCSsystem for the trial and the permanent installation periods. The spinalcord stimulator (SCS) 100 is an implantable device used to deliverelectrical pulse therapy to the spinal cord in order to treat chronicpain. The implantable components of the system consist of an ImplantablePulse Generator (IPG) 102 and a multitude of stimulation electrodes 130.The IPG 102 is implanted subcutaneously, no more than 30 mm deep in anarea that is comfortable for the patient while the stimulationelectrodes 130 are implanted directly in the epidural space. Theelectrodes 130 are wired to the IPG 102 via leads 140, 141 which keepthe stimulation pulses isolated from each other in order to deliver thecorrect therapy to each individual electrode 130.

The therapy delivered consists of electrical pulses with controlledcurrent amplitude ranging from +12.7 to −12.7 mA (current range 0-25.4mA). These pulses can be programmed in both length and frequency from 10μS to 2000 μS and 0.5 Hz to 1200 Hz. At any given moment, the sum of thecurrents sourced from the anodic electrodes 130 must equal the sum ofthe currents sunk by the cathodic electrodes 130. In addition, eachindividual pulse is bi-phasic, meaning that once the initial pulsefinishes another pulse of opposite amplitude is generated after a setholdoff period. The electrodes 130 may be grouped into stimulation setsin order to deliver the pulses over a wider area or to target specificareas, but the sum of the currents being sourced at any one given timemay not exceed 20 mA. A user can also program different stimulation sets(up to eight) with different parameters in order to target differentareas with different therapies.

FIG. 2 depicts an exploded view of an IPG 102. The IPG 102 consists oftwo major active components 104, 106, a battery 108, antenna 110, somesupport circuitry, and a multitude of output capacitors 112. The firstof the major active components is the microcontroller 104 transceiver104. It is responsible for receiving, decoding, and execution bothcommands and requests from the external remote. If necessary it passesthese commands or requests onto the second major component, the ASIC106. The ASIC 106 receives the digital data from the microcontroller 104and performs the entire signal processing to generate the signalsnecessary for stimulation. These signals are then passed onto thestimulation electrodes 130 in the epidural space.

The ASIC 106 is comprised of a digital section and an analog section.The digital section is divided into multiple sections including; TimingGenerators, Arbitration Control, Pulse Burst Conditioner, and ElectrodeLogic. The analog section receives the incoming pulses from the digitalsection and amplifies them in order to deliver the correct therapy.There are also a multitude of digital register memory elements that eachsection utilizes, both digital and analog.

The digital elements in the ASIC 106 are all made up of standard subsetsof digital logic including logic gates, timers, counters, registers,comparators, flip-flips, and decoders. These elements are ideal forprocessing the stimulation pulses as all of them can function extremelyfast-orders of magnitudes faster than the required pulse width. Theelements all function at one single voltage, usually 5.0, 3.3, 2.5, or1.8 volts.

The timing generators are the base of each of the stimulation sets. Itgenerates the actual rising and falling edge triggers for each phase ofthe bi-phasic pulse. It accomplishes this by taking the incoming clockthat is fed from the microcontroller 104 and feeding it into a counter.For the purpose of this discussion, assume the counter simply countsthese rising clock edges infinitely. The output of the counter is fedinto six different comparators. The comparators other input is connectedto specific registers that are programmed by the microcontroller 104.When the count equals the value stored in the register, the comparatorasserts a positive signal.

The first comparator is connected to the SET signal of a SR flip flop.The SR flip flop stays positive until the RESET signal is asserted,which the second comparator is connected to. The output of the SR flipflop is the first phase of the bi-phasic pulse. Its rising & fallingedges are values stored in the registers and programmed by themicrocontroller 104. The third and fourth comparators & registers workin exactly the same way to produce the second phase of the bi-phasicpulse using the second SR flip flop.

The fifth comparator is connected the RESET of the final SR-Flip flop inthe timing generator. This flip flop is SET by the first comparator,which is the rising edge of the first pulse. The RESET is then triggeredby the value the microprocessor programmed into the register connectedto the comparator. This allows for a ‘holdoff’ period after the fallingedge of the second pulse. The output of this third SR flip flop can bethought of as an envelope of the biphasic pulses indicating when thisparticular timing generator is active.

The final comparator of the system is once again connected to a registerthat stores the frequency values from the microprocessor. Essentiallywhen the count reaches this value it triggers the comparator which isfed back to the counter to reset it to zero and beginning the entirepulse generation cycle again. The ASIC 106 may contain many of thesetiming generators as each can control anywhere from two to all of theelectrodes 130 connected to the IPG 102 at a time. However, when thereis more than one timing generator and multiple channels have beenactively programmed then there needs to be a mechanism for suppressing asecond channel from turning on when another is already active.

The next circuit block contained in the IPG 102 is the arbitrator. Thearbitrator functions by looking at each of the timing generators'envelope signals and makes sure only one can be active at a time. If asecond tries to activate then the arbitrator suppresses that signal.

The arbitrator accomplishes this by bringing each of the channelenvelope signals into a rising edge detection circuit. Once one istriggered it is fed into the SET pin of an SR flip flop. The output ofthis SR-flip flop is fed into all of the other rising edge detectors inorder to suppress them from triggering. The channel envelope signal isalso fed into a falling-edge detector which is then fed into the RESETof the same SR flip flop. The output of the SR flip flops are thenconnected to switches whose outputs are all tied together that turnon/off that channels particular biphasic pulse train. Therefore, theoutput of this circuit element is a single bi-phasic pulse train and asignal designating which timing generator that particular pulse train issourced from. Essentially, the circuit looks for a channel to go active.Once it finds one it suppresses all others until that channel becomesinactive.

The next section of the circuit works very similarly to the timinggenerators to create a high speed burst pulse train that is thencombined with the stimulation pulse train to create a bursted bi-phasicpulse train if desired.

It accomplishes this by taking the incoming clock that is fed from themicrocontroller 104 and feeding it into a counter. The counter can countthese rising clock edges infinitely. The counter is only active during asingle phase of the bi-phasic signal and begins counting as soon as therising edge is detected. The output of the counter is fed into acomparator, along with a microcontroller-programmed register, whoseoutput is connected to the reset pin on the counter. Therefore, thiscounter will simply count to a programmed value and reset. Thisprogrammed value is the burst frequency.

The output of the comparator is then fed into an edge detection circuitand then a flip flop that combines it with the actual stimulation pulsetrain to create a single phase bursted stimulation pulse. The entirecircuit is duplicated for the second phase of the signal resulting inthe desired bursted bi-phasic pulse train. The stimulation signal is nowhanded over to the electrode logic stage.

The electrode logic conditions and directs the bi-phasic signals to theanalog section of the ASIC 106. At this point, the bi-phasic signalscontain all of the pertinent timing information, but none of therequired amplitude information. The incoming signals include thebi-phasic pulse train and another signal designating which timinggenerator the current active train came from. Each electrode logic cellhas a register for each timing generator that stores this particularelectrode's 130 amplitude values for that timing generator. Theelectrode logic cell uses the designation signal to determine whichregister to pull the amplitude values from, e.g. if the third timinggenerator is passed through the arbitration circuit then the electrodelogic would read the value from the third register.

Once the value is pulled from the register, it goes through a series oflogic gates. The gates first determine that the electrode 130 should beactive. If not, no further action is taken and the analog section of theelectrode output is not activated, thereby saving precious battery 108power. Next, a determination is made if the particular electrode 130 isan anode or cathode. If the electrode is deemed to be an anode, theelectrode logic passes the amplitude information and the biphasic signalto the positive current (digital to analog converter) DAC in the analogsection of the ASIC 106. If the electrode is deemed to be a cathode, theelectrode logic passes the amplitude information and the biphasic signalto the negative current DAC in the analog section of the ASIC 106. Theelectrode logic circuit must make these decisions for each phase of thebi-phasic signal as every electrode 130 will switch between being ananode and a cathode.

The analog elements in the ASIC 106 are uniquely designed in order toproduce the desired signals. The basis of analog IC design is the fieldeffect transistor (FET) and the type of high current multiple outputdesign required in SCS 100 means that the bulk of the silicon in theASIC 106 will be dedicated to the analog section.

The signals from the electrode output are fed into each current DAC whenthat specific electrode 130 should be activated. Each electrode 130 hasa positive and a negative current DAC, triggered by the electrode logicand both are never active at the same time. The job of each current DACis, when activated, to take the digital value representing a stimulationcurrent amplitude and produce an analog representation of this value tobe fed into the output stage. This circuit forms half of the barrierbetween the digital and analog sections of the ASIC 106.

The digital section of the ASIC 106 is built upon a technology that onlyallows small voltages to exist. In moving to the analog section, theoutput of the current DAC (which is a low level analog signal) must beamplified to a higher voltage for use in the analog section. The circuitthat performs this task is called a power level shifter. Because thiscircuit is built upon two different manufacturing technologies andrequires high precision analog circuits built upon a digital base, itcan be difficult to implement.

Once the voltages have been converted for usage in the analog portion ofthe ASIC 106 the voltages are passed on to the output current stages.There are two current sources per electrode output. One will source apositive current and one will sink a negative current, but both willnever be active simultaneously. The current sources themselves are madeup of analog elements similar to a Howland current source. There is aninput stage, and an amplification stage with feedback through a sensingcomponent to maintain the constant current. The input stage takes theanalog voltage values from the power level shifter and produces anoutput pulse designated for the amplifier. The amplifier then createsthe pulses of varying voltages but constant current flow. The sourcesare capable of sourcing or sinking up to 12.7 mA at 0.1 mA resolutioninto a load of up to 1.2 k Ohms. This translates into range of 15 volts,which will vary depending on the load in order to keep the currentconstant.

The microcontroller 104 to ASIC 106 interface is designed to be assimple as possible with minimal bus ‘chatter’ in order to save battery108 life. The ASIC 106 can be a collection of registers programmed via astandard I²C or SPI bus. Since the ASIC 106 is handling all the powermanagement, there will also be a power good (PG) line between the twochips 104, 106 in order to let the microcontroller 104 know when it issafe to power up. The ASIC 106 will also need to use a pin on themicrocontroller 104 in order to generate a hardware interrupt in caseanything goes awry in the ASIC 106. The final connection is the timebase for all of the stimulation circuitry. The ASIC 106 will require twoclocks, one for its internal digital circuitry which will be feddirectly from the microcontroller 104 clock output, and one to base allstimulation off of which will need to be synthesized by themicrocontroller 104 and fed to the ASIC 106. All commands and requeststo the ASIC 106 will be made over the I²C or SPI bus and will involvesimply reading a register address or writing to a register. Even whenthe ASIC 106 generates a hardware interrupt, it will be theresponsibility of the microcontroller 104 to poll the ASIC 106 anddetermine the cause of the interrupt.

The wireless interface is based upon the FCCs MedRadio standardoperating in the 402-405 MHz range utilizing up to 10 channels fortelemetry. The protocol implemented is chosen to minimize transmissionand maximize battery 108 life. All processing will take place on theuser remote/programmer and the only data transmitted is exactly whatwill be used in the microcontroller 104 to ASIC 106 bus. That is, all ofthe wireless packets will contain necessary overhead information alongwith only a register address, data to store in the register, and acommand byte instructing the microcontroller 104 what to do with thedata. The overhead section of the wireless protocol will containsynchronization bits, start bytes, an address which is synchronized withthe IPG's 102 serial number, and a CRC byte to assure propertransmission. The packet length is kept as small as possible in order tomaintain battery 108 life. Since the IPG 102 cannot listen for packetsall the time due to battery 108 life, it cycles on for a duty cycle ofless than 0.05% of the time. This time value can be kept small as longas the data packets are also small. The user commands needed to run thesystem are executed by the entire system using flows.

The IPG 102 uses an implantable grade Li ion battery 108 with 215 mAHrwith zero volt technology. The voltage of the battery 108 at fullcapacity is 4.1 V and it supplies current only until it is drained up to3.3 V which is considered as 100% discharged. The remaining capacity ofthe battery 108 can be estimated at any time by measuring the voltageacross the terminals. The maximum charge rate is 107.5 mA. A ConstantCurrent, Constant Voltage (CCCV) type of regulation can be applied forfaster charging of the battery 108.

The internal secondary coil 109 is made up of 30 turns of 30 AWG coppermagnet wires. The ID, OD, and the thickness of the coil are 30, 32, and2 mm, respectively. Inductance L2 is measured to be 58 uH, a 80 nFcapacitor is connected to it to make a series resonance tank at 74 kHzfrequency. In the art of induction charging, two types of rectifiers areconsidered to convert the induced AC into usable DC, either a bridgefull wave rectifier or a voltage doubler full wave rectifier. To obtaina higher voltage, the voltage double full wave rectifier is used in thisapplication. The rectifier is built with high speed Schottky diodes toimprove its function at high frequencies of the order 100 kHz. A Zenerdiode and also a 5V voltage regulator are used for regulation. Thiscircuit will be able to induce AC voltage, rectify to DC, regulate to 5Vand supply 100 mA current to power management IC that charges theinternal battery 108 by CCCV regulation.

The regulated 5V 100 mA output from the resonance tank is fed to, forexample, a Power Management Integrated Circuit (PMIC) MCP73843. Thisparticular chip was specially designed by Microchip to charge a Li ionbattery 108 to 4.1 V by CCCV regulation. The fast charge current can beregulated by changing a resistor; it is set to threshold current of 96mA in the example circuit. The chip charges the battery 108 to 4.1V aslong as the current received is more than 96 mA. However, if the supplycurrent drops below 96 mA, it stops to charge the battery 108 until thesupply is higher than 96 again. For various practical reasons, if thedistance between the coils increases, the internal secondary coil 109receives lesser current than the regulated value, and instead ofcharging the battery 108 slowly, it pauses the charging completely untilit receives more than 96 mA. It is understood to those with skill in theart that other power management chips can be used and the powermanagement chip is not limited to the PMIC MCP738432 chip.

All the functions of the IPG 102 are controlled from outside using ahand held remote controller specially designed for this device. Alongwith the remote control, an additional control is desirable to operatethe IPG 102 if the remote control was lost or damaged. For this purposea Hall effect based magnet switch was incorporated to either turn ON orturn OFF the IPG 102 using an external piece of magnet. Magnet switchacts as a master control for the IPG 102 to turn on or off. A south poleof sufficient strength turns the output on and a north pole ofsufficient strength turns the output off. The output is latched so thatthe switch continues to hold the state even after the magnet is removedfrom its vicinity.

The IPG 102 is an active medical implant that generates an electricalsignal that stimulates the spinal cord. The signal is carried through astimulation lead 140 that plugs directly into the IPG 102. The IPG 102recharges wirelessly through an induction coil 109, and communicates viaRF radio antenna 110 to change stimulation parameters. The IPG 102 isimplanted up to 3 cm below the surface of the skin and can be fixed tothe fascia by passing two sutures through holes in the epoxy header 114.The leads 140 are electrically connected to the IPG 102 through a leadcontact system 116, a cylindrical spring-based contact system withinter-contact silicone seals. The leads 140 are secured to the IPG 102with a set screw 117 that actuates within locking housing 118. Set screwcompression on the lead's 140 fixation contact can be governed by adisposable torque wrench. The wireless recharging is achieved byaligning the exterior induction coil on the charger with the internalinduction coil 109 within the IPG 102. The RF antenna within theremote's dongle 200 communicates with the RF antenna 110 in the IPG's102 epoxy header 114. FIG. 2 illustrates an exploded view of the IPG 102assembly.

The IPG 102 is an assembly of a hermetic titanium (6Al-4V) casing 120which houses the battery 108, circuitry 104, 106, and charging coil 109.The IPG 102 further includes an epoxy header 114 (see FIG. 6), whichhouses the lead contact assembly 116, locking housing 118, and RFantenna 110 (see FIGS. 6 and 7). The internal electronics are connectedto the components within the epoxy head through a hermetic feedthrough122, as shown in FIG. 3. The feedthrough 122 is a titanium (6Al-4V)flange with an alumina window and gold trimming. Within the aluminawindow are thirty-four platinum-iridium (90-10) pins that interfaceinternally with a direct solder to the circuit board, and externallywith a series of platinum iridium wires laser-welded to the antenna 110and lead contacts 126. The IPG 102 interfaces with 32 electricalcontacts 126, which are arranged in four rows of eight contacts 126.Thirty-two of the feedthrough's 122 pins 124 interface with the contacts126, while two interface with the antenna 110, one to the ground planeand one to the antenna 110 feed.

FIGS. 4 and 5 depict a lead contact system 115 and assembly 116,respectively. The lead contacts 126 consist of an MP35N housing 128 witha platinum-iridium 90-10 spring 129. Each contact 126 is separated by asilicone seal 127. At the proximal end of each stack of 8 contacts 126is a titanium (6Al-4V) cap 125 which acts as a stop for the lead 140. Atthe distal end is a titanium (6Al-4V) set screw 119 and block 118 forlead fixation. At the lead entrance point is a silicone tube 123 whichprovides strain relief as the lead 140 exits the head unit 114, andabove the set screw 119 another silicone tube 131 with a small internalcanal allows the torque wrench to enter but does not allow the set screw119 to back out. In addition to the contacts 126 and antenna 110, theheader 114 also contains a radiopaque titanium (6Al-4V) tag 132 whichallows for identification of the device under fluoroscopy. The overmoldof the header 114 is Epotek 301, a two-part, biocompatible epoxy. FIGS.4, 5, 6, and 7 depict illustrations of lead contact system 115, leadcontact assembly 116, head unit assembly 114, and RF antenna 110,respectively.

Internal to the titanium (6Al-4V) case 120 are the circuit board 105,battery 108, charging coil 109, and internal plastic support frame. Thecircuit board 105 can be a multi-layered FR-4 board with copper tracesand solder mask coating. Non-solder masked areas of the board can beelectroless nickel immersion gold. The implantable battery 108, allsurface mount components, ASIC 106, microcontroller 104, charging coil109, and feedthrough 122 will be soldered to the circuit board 105. Theplastic frame, made of either polycarbonate or ABS, will maintain thebattery's 108 position and provide a snug fit between the circuitry 105and case 120 to prevent movement. The charging coil 109 is a woundcoated copper.

Leads

The percutaneous stimulation leads 140, as depicted in FIG. 8, are afully implantable electrical medical accessory to be used in conjunctionwith the implantable SCS 100. The primary function of the lead is tocarry electrical signals from the IPG 102 to the target stimulation areaon the spinal cord. Percutaneous stimulation leads 140 providecircumferential stimulation. The percutaneous stimulation leads 140provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and stimulation area. The leads 140 are surgicallyimplanted through a spinal needle, or epidural needle, and are driventhrough the spinal canal using a steering stylet that passes through thecenter of the lead 140. The leads 140 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 140. The leads 140 are secured at theproximal end with a set-screw 119 on the IPG 102 which applies radialpressure to a blank contact on the distal end of the proximal contacts.

The percutaneous stimulation leads 140 consist of a combination ofimplantable materials. Stimulation electrodes 130 at the distal end andelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy. This alloy is utilized for its bio-compatibilityand electrical conductivity. The electrodes 130 are geometricallycylindrical. The polymeric body of the lead 140 is polyurethane, chosenfor its bio-compatibility, flexibility, and high lubricity to decreasefriction while being passed through tissue. The polyurethane tubing hasa multi-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. The wires 146are laser welded to the contacts and electrodes 130, creating anelectrical connection between respective contacts on the proximal anddistal ends. The leads 140 employ a platinum-iridium plug 148, moldedinto the distal tip of the center lumen 142 to prevent the tip of thesteering stylet from puncturing the distal tip of the lead 140. Leads140 are available in a variety of 4 and 8 electrode 130 configurations.These leads 140 have 4 and 8 proximal contacts (+1 fixation contact),respectively. Configurations vary by electrode 130 number, electrode 130spacing, electrode 130 length, and overall lead 140 length.

The paddle stimulation leads 141, as depicted in FIG. 9, are a fullyimplantable electrical medical accessory to be used in conjunction withthe implantable SCS 100. The primary function of the paddle lead 141 isto carry electrical signals from the IPG 102 to the target stimulationarea on the spinal cord. The paddle leads 141 provide uni-directionstimulation across a 2-dimensional array of electrodes 130, allowing forgreater precision in targeting stimulation zones. The paddle stimulationleads 141 provide a robust, flexible, and bio-compatible electricconnection between the IPG 102 and stimulation area. The leads 141 aresurgically implanted through a small incision, usually in conjunctionwith a laminotomy or laminectomy, and are positioned using forceps or asimilar surgical tool. The leads 141 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 141. The leads 141 are secured at theproximal end with a set-screw on the IPG 102 which applies radialpressure to a fixation contact on the distal end of the proximalcontacts.

The paddle stimulation leads 141 consist of a combination of implantablematerials. Stimulation electrodes 130 at the distal end and electricalcontacts at the proximal end are made of a 90-10 platinum-iridium alloyutilized for its bio-compatibility and electrical conductivity. Thepolymeric body of the lead 141 is polyurethane, chosen for itsbio-compatibility, flexibility, and high lubricity to decrease frictionwhile being passed through tissue. The polyurethane tubing has amulti-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. At the distaltip of the paddle leads 141 is a 2-dimensional array of flat rectangularelectrodes 130 molded into a flat silicone body 149. In an embodiment,one side of the rectangular electrodes 130 is exposed, providinguni-directional stimulation. The wires 146 are laser welded to thecontacts and electrodes 130, creating an electrical connection betweenrespective contacts on the proximal and distal ends. Also molded intothe distal silicone paddle is a polyester mesh 147 adding stability tothe molded body 149 while improving aesthetics by covering wire 146routing. The number of individual 8-contact leads 141 used for eachpaddle 141 is governed by the number of electrodes 130. Electrodes 130per paddle 141 range from 8 to 32, split into between one and fourproximal lead 141 ends. Each proximal lead 141 has 8 contacts (+1fixation contact). Configurations vary by electrode 130 number,electrode 130 spacing, electrode length, and overall lead length.

The lead extensions 150, as depicted in FIG. 10, are a fully implantableelectrical medical accessory to be used in conjunction with theimplantable SCS 100 and either percutaneous 140 or paddle 141 leads. Theprimary function of the lead extension 150 is to increase the overalllength of the lead 140, 141 by carrying electrical signals from the IPG102 to the proximal end of the stimulation lead 140, 141. This extendsthe overall range of the lead 140, 141 in cases where the length of theprovided leads 140, 141 are insufficient. The lead extensions 150provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and proximal end of the stimulation lead 140, 141.The extensions 150 may be secured mechanically to the patient usingeither an anchor or a suture passed through tissue and tied around thebody of the extension 150. Extensions 150 are secured at the proximalend with a set-screw 119 on the IPG 102 which applies radial pressure toa fixation contact on the distal end of the proximal contacts of theextension 150. The stimulation lead 140, 141 is secured to the extension150 in a similar fashion, using a set screw 152 inside the molded tip ofextension 150 to apply a radial pressure to the fixation contact at theproximal end of the stimulation lead 140, 141.

The lead extension 150 consists of a combination of implantablematerials. At the distal tip of the extension 150 is a 1×8 array ofimplantable electrical contacts 154, each consisting of MP35 housing 128and 90-10 platinum-iridium spring. A silicone seal 127 separates each ofthe housings 128. At the proximal end of the contacts is a titanium(6Al4V) cap which acts as a stop for the lead, and at the distal tip, atitanium (6Al4V) block and set screw 152 for lead fixation. Theelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy utilized for its bio-compatibility and electricalconductivity. The polymeric body 156 of the lead 150 is polyurethane,chosen for its bio-compatibility, flexibility, and high lubricity todecrease friction while being passed through tissue. The polyurethanetubing 158 has a multi-lumen cross section, with one center lumen 142and eight outer lumens 144. The center lumen 142 acts as a canal tocontain the steering stylet during implantation, while the outer lumens144 provide electrical and mechanical separation between the wires 146that carry stimulation from the proximal contacts to distal electrodes.These wires 146 are a bundle of MP35N strands with a 28% silver core.The wires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. Each leadextension 150 has 8 proximal cylindrical contacts (+1 fixation contact).

The lead splitter 160, as depicted in FIG. 11, is a fully implantableelectrical medical accessory which is used in conjunction with the SCS100 and typically a pair of 4-contact percutaneous leads 140. Theprimary function of the lead splitter 160 is to split a single lead 140of eight contacts into a pair of 4 contact leads 140. The splitter 160carries electrical signals from the IPG 102 to the proximal end of two4-contact percutaneous stimulation leads 140. This allows the surgeonaccess to more stimulation areas by increasing the number of stimulationleads 140 available. The lead splitter 160 provides a robust, flexible,and bio-compatible electrical connection between the IPG 102 andproximal ends of the stimulation leads 140. The splitters 160 may besecured mechanically to the patient using either an anchor or a suturepassed through tissue and tied around the body of the splitter 160.Splitters 160 are secured at the proximal end with a set-screw 119 onthe IPG 102 which applies radial pressure to a fixation contact on thedistal end of the proximal contacts of the splitter 160. The stimulationleads 140 are secured to the splitter 160 in a similar fashion, using apair of set screws inside the molded tip of splitter 160 to apply aradial pressure to the fixation contact at the proximal end of eachstimulation lead 140.

The lead splitter 160 consists of a combination of implantablematerials. At the distal tip of the splitter 160 is a 2×4 array ofimplantable electrical contacts 162, with each contact 162 consisting ofMP35 housing 128 and 90-10 platinum-iridium spring. A silicone seal 127separates each of the housings 128. At the proximal end of each row ofcontacts 162 is a titanium (6Al4V) cap which acts as a stop for thelead, and at the distal tip, a titanium (6Al4V) block and set screw forlead fixation. The electrical contacts at the proximal end of thesplitter 160 are made of a 90-10 platinum-iridium alloy utilized for itsbio-compatibility and electrical conductivity. The polymeric body 164 ofthe lead 160 is polyurethane, chosen for its bio-compatibility,flexibility, and high lubricity to decrease friction while being passedthrough tissue. The polyurethane tubing 166 has a multi-lumen crosssection, with one center lumen 142 and eight outer lumens 144. Thecenter lumen 142 acts as a canal to contain the steering stylet duringimplantation, while the outer lumens 144 provide electrical andmechanical separation between the wires 146 that carry stimulation fromthe proximal contacts to distal electrodes 130. These wires 146 are abundle of MP35N strands with a 28% silver core. The wires 146 areindividually coated with ethylene tetrafluoroethylene (ETFE), to providean additional non-conductive barrier. Each lead splitter 160 has 8proximal contacts (+1 fixation contact), and 2 rows of 4 contacts 162 atthe distal end.

Anchors

The lead anchor 170, as depicted in FIGS. 12 and 13, is a fullyimplantable electrical medical accessory which is used in conjunctionwith both percutaneous 140 and paddle 141 stimulation leads. The primaryfunction of the lead anchor 170 is to prevent migration of the distaltip of the lead 140, 141 by mechanically locking the lead 140, 141 tothe tissue. There are currently two types of anchors 170, a simplesleeve 171, depicted in FIG. 12, and a locking mechanism 172, depictedin FIG. 13, and each has a slightly different interface. For the simplesleeve type anchor 171, the lead 140, 141 is passed through the centerthru-hole 174 of the anchor 171, and then a suture is passed around theoutside of the anchor 171 and tightened to secure the lead 140, 141within the anchor 171. The anchor 171 can then be sutured to the fascia.The locking anchor 172 uses a set screw 176 for locking purposes, and abi-directional disposable torque wrench for locking and unlocking.Tactile and audible feedback is provided for both locking and unlocking.

Both anchors 171, 172 can be molded from implant-grade silicone, but thelocking anchor 172 uses an internal titanium assembly for locking. The3-part mechanism is made of a housing 175, a locking set screw 176, anda blocking set screw 177 to prevent the locking set screw from back out.All three components can be titanium (6Al4V). The bi-directional torquewrench can have a plastic body and stainless steel hex shaft.

Wireless Dongle

The wireless dongle 200 is the hardware connection to asmartphone/mobile 202 or tablet 204 that allows communication betweenthe trial generator 107 or IPG 102 and the smartphone/mobile device 202or tablet 204, as illustrated in FIG. 14. During the trial or permanentimplant phases, the wireless dongle 200 is connected to the tablet 204through the tablet 204 specific connection pins and the clinicianprogrammer software on the tablet 204 is used to control the stimulationparameters. The commands from the clinician programmer software aretransferred to the wireless dongle 200 which is then transferred fromthe wireless dongle 200 using RF signals to the trial generator 107 orthe IPG 102. Once the parameters on the clinician programmers have beenset, the parameters are saved on the tablet 204 and can be transferredto the patient programmer software on the smartphone/mobile device 202.The wireless dongle 200 is composed of an antenna, a microcontroller(having the same specifications as the IPG 102 and trial generator 107),and a pin connector to connect with the smartphone/mobile device 202 andthe tablet 204.

Charger

The IPG 102 has a rechargeable lithium ion battery 108 to power itsactivities. An external induction type charger 210 (FIG. 1) wirelesslyrecharges the included battery 108 inside the IPG 102. The charger 210is packaged into a housing and consists of a rechargeable battery, aprimary coil of wire and a printed circuit board (PCB) containing theelectronics. In operation, charger 210 produces a magnetic field andinduces voltage into the secondary coil 109 in the IPG 102. The inducedvoltage is then rectified and used to charge the battery 108 inside theIPG 102. To maximize the coupling between the coils, both internal andexternal coils are combined with capacitors to make them resonate at aparticular common frequency. The coil acting as an inductor L forms anLC resonance tank. The charger uses a Class-E amplifier topology toproduce the alternating current in the primary coil around the resonantfrequency. The charger 210 features include, but are not limited to:

Charge IPG 102 wirelessly

Charge up to a maximum depth of 30 mm

Integrated alignment sensor indicates alignment between the charger andIPG 102 resulting in higher power transfer efficiency

Alignment sensor provides audible and visual feedback to the user

Compact and Portable

A protected type of cylindrical Li ion battery is used as the charger210 battery. A Class-E power amplifier topology is a much used type ofamplifier for induction chargers, especially for implantable electronicmedical devices. Due to the Class-E power amplifier's relatively hightheoretical efficiency it is often used for devices where highefficiency power transfer is necessary. A 0.1 ohm high wattage resistoris used in series to sense the current through this circuit.

The primary coil L1 is made by 60 turns of Litz wire type 100/44-100strands of 44 AWG each. The Litz wire solves the problem of skin effectand keeps its impedance low at high frequencies. Inductance of this coilwas initially set at 181 uH, but backing it with a Ferrite plateincreases the inductance to 229.7 uH. The attached ferrite plate focusesthe produced magnetic field towards the direction of the implant. Such asetup helps the secondary coil receive more magnetic fields and aids itto induce higher power.

When the switch is ON, the resonance is at frequency

$f = \frac{1}{2\pi \sqrt{L\; 1C\; 2}}$

When the switch is OFF, it shifts to

$f = \frac{1}{2\pi \sqrt{{L1}\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

In a continuous operation the resonance frequency will be in the range

$\frac{1}{2\pi \sqrt{L\; 1C\; 2}} < f < \frac{1}{2\pi \sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

To make the ON and OFF resonance frequencies closer, a relatively largervalue of C1 can be chosen by a simple criteria as follows

C1=nC2; a value of n=4 was used in the example above; in most cases3<n<10.

The voltages in these Class-E amplifiers typically go up to the order of300VAC. Capacitors selected must be able to withstand these highvoltages, sustain high currents and still maintain low Effective SeriesResistance (ESR). Higher ESRs result in unnecessary power losses in theform of heat. The circuit is connected to the battery through aninductor which acts as a choke. The choke helps to smoothen the supplyto the circuit. The N Channel MOSFET acts as a switch in this Class-Epower amplifier. A FET with low ON resistance and with high draincurrent Id is desirable.

In summary, the circuit is able to recharge the IPG 102 battery 108 from0 to 100% in approximately two hours forty-five minutes with distancebetween the coils being 29 mm. The primary coil and the Class-Eamplifier draws DC current of 0.866 A to achieve this task. To improvethe efficiency of the circuit, a feedback closed loop control isimplemented to reduce the losses. The losses are minimum when the MOSFETis switched ON and when the voltage on its drain side is close to zero.

The controller takes the outputs from operational amplifiers, checks ifthe outputs meet the criteria, then triggers the driver to switch ON theMOSFET for the next cycle. The controller can use a delay timer, an ORgate and a 555 timer in monostable configuration to condition the signalfor driver. When the device is switched ON, the circuit will notfunction right away as there is no active feedback loop. The feedbackbecomes active when the circuit starts to function. To provide an activefeedback loop, an initial external trigger is applied to jump start thesystem.

Alignment Sensor

The efficiency of the power transfer between the external charger 210and the internal IPG 102 will be maximum only when the charger 210 andIPG 102 are properly aligned. An alignment sensor is provided to ensureproper alignment as part of the external circuit design and is based onthe principle of reflected impedance. When the external coil is broughtcloser to the internal coil, the impedance of both circuits change. Thesensing is based on measuring the reflected impedance and testingwhether it crosses the threshold. A beeper provides an audible feedbackto the patient and a LED provides visual feedback.

When the impedance of the circuit changes, the current passing throughit also changes. A high power 0.1 ohm resistor can be used in the seriesof the circuit to monitor the change in current. The voltage drop acrossthe resistor is amplified 40 times and then compared to a fixedthreshold value using an operational amplifier voltage comparator. Theoutput is fed to a timer chip which in turn activates the beeper and LEDto provide feedback to the user.

The circuit can sense the alignment up to a distance of approximately 30mm. The current fluctuation in the circuit depends on more factors thanreflected impedance alone and the circuit is sensitive to otherparameters of the circuit as well. To reduce the sensitivity related toother parameters, one option is to eliminate interference of all theother factors and improve the functionality of the reflected impedancesensor—which is very challenging to implement within the limited spaceavailable for circuitry. Another option is to use a dedicated sensorchip to measure the reflected impedance.

A second design uses sensors designed for proximity detector or metaldetectors for alignment sensing. Chips designed to detect metal bodiesby the effect of Eddy currents on the HF losses of a coil can be usedfor this application. The TDE0160 is an example of such a chip.

The external charger is designed to work at 75 to 80 kHz, whereas theproximity sensor was designed for 1 MHz. The sensor circuit is designedto be compatible with the rest of the external and is fine tuned todetect the internal IPG 102 from a distance of approximately 30 mm.

Programmer

The Clinician Programmer is an application that is installed on a tablet204. It is used by the clinician to set the stimulation parameters onthe trial generator 107 or IPG 102 during trial and permanentimplantation in the operating room. The clinician programmer is capableof saving multiple settings for multiple patients and can be used toadjust the stimulation parameters outside of the operations room. It iscapable of changing the stimulation parameters though the RF wirelessdongle 200 when the trial generator 107 or IPG 102 which has beenimplanted in the patient is within the RF range. In addition, it is alsocapable of setting or changing the stimulation parameters on the trialgenerator 107 and/or the IPG 102 through the internet when both thetablet 204 and the Patient Programmers on a smartphone/mobile device 202both have access to the internet.

The Patient Programmer is an application that is installed on asmartphone/mobile device 202. It is used by the patient to set thestimulation parameters on the trial generator 107 or IPG 102 after trialand permanent implantation outside the operating room. The clinicianprogrammer is capable of saving multiple settings for multiple patientsand can be transferred to the Patient Programmer wirelessly when theClinician Programmer tablet 204 and the Patient Programmersmartphone/mobile device 202 are within wireless range such as Bluetoothfrom each other. In the scenario where the Clinician Programmer tablet204 and the Patient Programmer smartphone/mobile device 202 are out ofwireless range from each other, the data can be transferred through theinternet where both devices 202, 204 have wireless access such as Wi-Fi.The Patient Programmer is capable of changing the stimulation parameterson the trial generator 107 or IPG 102 though the RF wireless dongle 200when the trial generator 107 or IPG implanted in the patient is withinthe RF range.

Tuohy Needle

The Tuohy needle 240, as depicted in FIG. 15, is used in conjunctionwith a saline-loaded syringe for loss-of-resistance needle placement,and percutaneous stimulation leads 140, for lead 140 placement into thespinal canal. The Tuohy epidural needle 240 is inserted slowly into thespinal canal using a loss-of-resistance technique to gauge needle 240depth. Once inserted to the appropriate depth, the percutaneousstimulation lead 140 is passed through the needle 240 and into thespinal canal.

The epidural needle 240 is a non-coring 14G stainless steel spinalneedle 240 and will be available in lengths of 5″ (127 mm) and 6″(152.4). The distal tip 242 of the needle 240 has a slight curve todirect the stimulation lead 140 into the spinal canal. The proximal end246 is a standard Leur-Lock connection 248.

Stylet

The stylet 250, as depicted in FIG. 16, is used to drive the tip of apercutaneous stimulation lead 140 to the desired stimulation zone byadding rigidity and steerability. The stylet 250 wire 252 passes throughthe center lumen 142 of the percutaneous lead 140 and stops at theblocking plug at the distal tip of the lead 140. The tip of the stylet250 comes with both straight and curved tips. A small handle 254 is usedat the proximal end of the stylet 250 to rotate the stylet 250 withinthe center lumen 142 to assist with driving. This handle 254 can beremoved and reattached allowing anchors 170 to pass over the lead 140while the stylet 250 is still in place. The stylet 250 wire 252 is aPTFE coated stainless steel wire and the handle 254 is plastic.

Passing Elevator

The passing elevator 260, as depicted in FIG. 17, is used prior topaddle lead 141 placement to clear out tissue in the spinal canal andhelp the surgeon size the lead to the anatomy. The passing elevator 260provides a flexible paddle-shaped tip 262 to clear the spinal canal ofobstructions. The flexible tip is attached to a surgical handle 264.

The passing elevator 260 is a one-piece disposable plastic instrumentmade of a flexible high strength material with high lubricity. Theflexibility allows the instrument to easily conform to the angle of thespinal canal and the lubricity allows the instrument to easily passthrough tissue.

Tunneling Tool

The tunneling tool 270, as depicted in FIG. 18, is used to provide asubcutaneous canal to pass stimulation leads 140 from the entrance pointinto the spinal canal to the IPG implantation site. The tunneling tool270 is a long skewer-shaped tool with a ringlet handle 272 at theproximal end 274. The tool 270 is covered by a plastic sheath 276 with atapered tip 278 which allows the tool 270 to easily pass through tissue.Once the IPG 102 implantation zone is bridge to the lead 140 entrancepoint into the spinal canal, the inner core 275 is removed, leaving thesheath 276 behind. The leads 140 can then be passed through the sheath276 to the IPG 102 implantation site. The tunneling tool 270 is oftenbent to assist in steering through the tissue.

The tunneling tool 270 is made of a 304 stainless steel core with afluorinated ethylene propylene (FEP) sheath 276. The 304 stainless steelis used for its strength and ductility during bending, and the sheath276 is used for its strength and lubricity.

Torque Wrench

The torque wrench 280, as depicted in FIG. 19, is used in conjunctionwith the IPG 102, lead extension 150 and lead splitter 160 to tightenthe internal set screw 119, which provides a radial force against thefixation contact of the stimulation leads 140, 141, preventing the leads140, 141 from detaching. The torque wrench 280 is also used to lock andunlock the anchor 170. The torque wrench 280 is a small, disposable,medical instrument that is used in every SCS 100 case. The torque wrench280 provides audible and tactile feedback to the surgeon that the lead140, 141 is secured to the IPG 102, extension 150, or splitter 160, orthat the anchor 170 is in the locked or unlocked position.

The torque wrench 280 is a 0.9 mm stainless steel hex shaft 282assembled with a plastic body 284. The wrench's 280 torque rating isbi-directional, primarily to provide feedback that the anchor 170 iseither locked or unlocked. The torque rating allows firm fixation of theset screws 119, 152 against the stimulation leads 140, 141 withoutover-tightening.

Trial Patch

The trial patch is used in conjunction with the trialing pulse generator107 to provide a clean, ergonomic protective cover of the stimulationlead 140, 141 entrance point in the spinal canal. The patch is alsointended to cover and contain the trial generator 107. The patch is alarge, adhesive bandage that is applied to the patient post-operativelyduring the trialing stage. The patch completely covers the leads 140,141 and generator 107, and fixates to the patient with anti-microbialadhesive.

The patch is a watertight, 150 mm×250 mm anti-microbial adhesive patch.The watertight patch allows patients to shower during the trialingperiod, and the anti-microbial adhesive decreases the risk of infection.The patch will be made of polyethylene, silicone, urethane, acrylate,and rayon.

Magnetic Switch

The magnetic switch is a magnet the size of a coin that, when placednear the IPG 102, can switch it on or off. The direction the magnet isfacing the IPG 102 determines if the magnetic switch is switching theIPG 102 on or off.

FIG. 20 is a functional block diagram of a wireless charger 210according to the present invention. As discussed above, the wirelesscharger 210 has a micro usb port 4 which is connectable to an externalpower supply (not shown) to receive DC power for charging the lithiumion rechargeable battery 6. In the embodiment shown, the rechargeablebattery 6 is a 4.2V battery. A power management circuit 8 regulates thepower from the port 4 to proper voltage and current which is used tocharge the battery 6.

A processor 12 such as a microcontroller controls the charging processand is powered by the battery 6. Since the processor 12 uses 3.3V, avoltage regulator 10 connected to the battery 6 regulates the batteryvoltage down to 3.3V for powering the processor.

A vibrator such as a vibration motor 14 for producing a vibratingtactile feedback as well as a buzzer/speaker 16 for creating soundfeedback for the user are connected to and are controlled by theprocessor 12. The vibrator 12 is similar to those used for cellulartelephones, game controllers, tablets and the like. For example, avibration motor part number 28821 from Parallex Inc. of Rocklin, Calif.can be used.

A user interface 18 is connected to the processor 12 to interact withthe charger 210. The user interface 18 may include buttons and switchesfor turning the charger 210 on and off and for changing the volume ofthe sound and motor from the vibrator 14 and speaker 16.

A frequency generator 20 coupled to the processor 12 generates a highfrequency signal under the control of the processor. For example, in oneembodiment, the processor 12 controls the frequency generator 20 togenerate a high frequency charging signal of between 80 kHz and 90 kHz.A power amplifier 22 coupled to the frequency generator 20 generates anamplified charging signal under the control of the charging signal. Theamplifier 22 then feeds the amplified charging signal to a charging coil24, which is placed on one side of a printed circuit board (PCB). In oneembodiment, the amplified charging signal is a sine wave signal having apeak to peak voltage of 500V. The charging coil 24 generates a magneticfield for inducing power into the IPG 102 as will be discussed in moredetail herein.

In the IPG 102 as partially shown in FIG. 21, the receiving coil 109 isinductively and wirelessly coupled to the charging coil 24 when they arepositioned near each other. The magnetic field from the charging coil 24induces voltage in the receiving coil 109. As an example, the inducedvoltage is an oscillating voltage with a swing of +3V to −3V for a peakto peak voltage of 6V. A resonance tank 28 which includes a capacitorconnected in series with the receiving coil 109 comprises a resonancecircuit whose resonance frequency is tuned to the frequency of themagnetic field emanating from the charging coil 24. The induced voltageis rectified by a rectifier 30 to convert an oscillating voltage into aDC voltage. In the embodiment shown, the rectifier 30 is a full wavevoltage doubler rectifier so as to generate a 6V DC at its output. TheIPG processor 104 such as a microcontroller controls a switch 32connected across the rectifier 30 in parallel. One end of the switch 32is connected to a power management circuit 34 while the other end isconnected to ground Vss. In one embodiment, the switch 32 is a MOSFETtransistor that can be turned on or off by the processor 104. Normally,the switch 32 is turned off including the time when the battery 108 isbeing charged.

The power management circuit 34 receives the rectified DC voltage fromthe rectifier 30 and charges the rechargeable battery 108. Othercircuits 35 control the actual generation and controlling of the spinalcord stimulation signals.

In one embodiment, the amplifier 22 is a class-E amplifier as shown inFIG. 22. The amplifier 22 includes a switch 40 receiving the chargingsignal from the frequency generator 20, LC circuit including capacitors42,44 and a choke 37 connected to the voltage source 6. The choke 37 inthe embodiment shown is an inductor that helps to smooth the powersupply to the power amplifier circuit. The switch 40 as shown is anN-channel MOSFET which is controlled by the frequency generator 20. Insome operating conditions for this type of amplifier, it is possiblethat the switch 40 can connect the voltage source directly to ground,effectively creating an electrical short, however fleeting it may be.That creates at least two issues. First, there will be a rapid drain ofhigh current from source to ground which is a waste of power. Second,the short drops the source voltage significantly to cause a malfunctionin other parts of the charger 210. This may cause some circuits tobehave erratically which is highly undesirable.

One solution is to use a current limiting resistor in the current pathof the amplifier 22. Although that solution reduces the maximum draincurrent, it also reduces the current that goes into the charging coil24. A preferred solution is to use a current limiter 36 connectedbetween the power source 6 and the RF choke 37 of the amplifier 22 tolimit the current being provided to the threshold value. In oneembodiment, the current limiter 36 is an integrated circuit chipNCP380LSNAJAAT1G from ON Semiconductor of Phoenix, Ariz. In theembodiment shown, the current limiter 36 has been programmed to limitthe current to a maximum threshold current of 0.5 Amps. In the case of ashort circuit between the voltage source 6 and ground through the MOSFETswitch 40, the current limiter 36 will not allow the amplifier 22 todrain more than the maximum current set limit of 0.5 Amps, for example.The current limiter 36 also avoids a significant voltage drop of thevoltage source, thereby allowing the rest of the electronic circuits tofunction normally.

The wireless charger 210 produces a high frequency, high voltagemagnetic field. The charging coil 24 has a resistance and the resistivelosses will be dissipated in the form of heat. Heat raises thetemperature of the charger 210. When the temperature of the charger 210rises significantly, the charger might cause minor discomfort in most ofthe cases and minor tissue burns in some rare cases.

According to Standards IEC60I01-1, temperature of the surface of adevice that is in physical contact with a patient's body shall belimited to 41 C. To control the heat, a temperature sensor 26 is used tomonitor the temperature of the charger 210. The charging coil 24 is aflat round shaped coil on one side of the PCB. In one embodiment, aferrite plate can be disposed between the coil and the PCB. On the otherside of the PCB, the temperature sensor 26 such as a thermistor 26 isplaced behind the coil 24 and is coupled to the processor 12. In oneembodiment, the coil 24 has an inner diameter of about 25 mm and anouter diameter of about 60 mm. In one embodiment, the temperature sensor26 is placed in the middle of the coil at about 42.5 mm from the coilcenter between the inner winding and outer winding. In this way, thehighest temperature of the coil 24 can generally be measured.

The processor 12 is programmed to monitor the temperature from thetemperature sensor 26 regularly and when the monitored temperature risesabove a first threshold value, the processor turns off the frequencygenerator 20 and power amplifier 22. The processor 12 continues tomonitor the temperature from the temperature sensor 26 and will turn thefrequency generator 20 and power amplifier 22 back on automatically whenthe monitored temperature falls below a second threshold value. Thus,there is a hysteresis band between the high and low thresholds to avoidthe charger 210 from rapidly switching on and off near the settemperature limit. Turning off the charging function when necessary willreduce patient's discomfort and likely avoid any tissue burns. Oneexample of first and second threshold temperatures may be 40 C and 38 C,respectively.

According to another aspect of the present invention, a chargeralignment feature for more efficiently transferring power into the IPG102 will now be explained. A charge alignment software is stored in aninternal memory 13 of the microcontroller 12 and is executed when thecharger 210 is turned on. For this feature, a reflected impedance sensor38 is used to measure a reflected impedance to detect a reflectedimpedance of the charging coil 24. Thus, a charge alignment circuitcomprises the stored charge alignment software, processor 12 andreflected impedance sensor 38.

As shown in FIG. 23, the sensor 38 includes a small transformer 46having a primary coil connected to the charging coil 24 in series todetect the current flowing through the charging coil. The secondary coilof the transformer 46 is electromagnetically coupled to the primarycoil. The primary coil acts as a sensor to sense the voltage across thecharging coil. The sensor 38 also includes a rectifier 48 (e.g., halfwave rectifier in the embodiment shown) to rectify AC current from thesecondary coil of the transformer 46 into DC. In the embodiment shown,the DC voltage can swing between zero and about 5 Volt. The rectifiedvoltage represents a voltage level of the charging coil 24. A one-to-onevoltage divider 50 divides the DC voltage to make the DC voltagecompatible with the operating voltage (e.g., 3.3 Volt) of themicrocontroller 12. Zener diode 52 connected between the voltage divider50 and ground ensures that the voltage from the voltage divider 50 doesnot rise above the operating voltage of the processor 12 by sinkingcurrent to ground if it does.

The output of the reflected impedance sensor 38 is connected to theprocessor 12. Typically, when the IPG 102 is far away and not receivingany current from the charger 210, the output voltage of the sensor 38 isaround 2.3-2.4 Volt. On the other hand, when the charging coil 24 of thecharger 210 is perfectly aligned with the receiving coil 109 of the IPG102, i.e., the charging coil is directly above the receiving coil 109,the output voltage of the sensor 38 drops to around 1.6-1.7 Volt. Athreshold voltage value of around 1.8 Volt is set such that any outputvoltage of the reflected impedance sensor 38 below that thresholdvoltage value is considered to be aligned for maximum charging currenttransfer from the charger 210 to the IPG 102.

Under the control of the charge alignment software, the microcontroller12 continuously monitors and compares the output of the reflectedimpedance detected by the sensor 38 against the threshold value andcontrols the vibrator 14 and speaker 16 to provide audible and tactilefeedback to the user/patient based on the detected reflected impedancevalues. Thus, the outputs of the vibrator 14 and speaker 16 areindicative of the alignment of the charging coil 24 to the receivingcoil 109.

In one embodiment, as the sensor 38 output decreases towards thethreshold value, meaning that the charging coil 24 is becoming morealigned with the receiving coil 109, the processor 12 controls thevibrator 14 to vibrate at a lower rate. When the sensor 38 outputreaches and goes past the threshold value, the processor 12 stops thevibrator 14 from vibrating. For example, the initial vibrator frequencycan be 8-9 Hertz, vibrating for 0.1 second each time. As the charger 210comes closer to the IPG 102, the vibrating frequency can correspondinglydecrease to 1-2 Hertz with the same 0.1 second vibrating duration. Whenthe charger 210 is fully aligned, i.e., the output of the sensor 38 hasreached the threshold value, then the processor 12 stops the vibrator 14from vibrating to indicate that the charger 210 is now fully alignedwith the IPG 102. Thus, in this embodiment, although the vibration ratedecreases, vibration is continuous until the charging coil 24 is fullyaligned with the receiving coil 109.

At the same time, the processor 12 controls the speaker 16 to generate atone (e.g., beeps) having the same frequency and duration as thevibrator 14. In other words, the processor 12 can control the speaker 16to make a tone at the initial interval of 8-9 Hertz, generating thesound for 0.1 second each time. As the charger 210 comes closer to theIPG 102, the tone can correspondingly decrease to 1-2 Hertz with thesame 0.1 second sound duration. When the charger 210 is fully aligned,i.e., the output of the sensor 38 has reached the threshold value, thenthe processor 12 stops the speaker 16 from generating any sound toindicate that the charger 210 is now fully aligned with the IPG 102.

In another embodiment, as the sensor 38 output decreases towards thethreshold value, the processor 12 controls the vibrator 14 to vibrate ata higher rate. When the sensor 38 output reaches the threshold value,the processor 12 controls the vibrator 14 to vibrate constantly. Forexample, the initial vibrator frequency can be 1 Hertz, vibrating for0.1 second each time. As the charger 210 comes closer to the IPG 102,the vibrating frequency can correspondingly increase to 8-9 Hertz withthe same 0.1 second vibrating duration. When the charger 210 is fullyaligned, i.e., the output of the sensor 38 has reached the thresholdvalue, then the vibration can be constant. Thus, in this embodiment,although the vibration rate increases, vibration is continuous until thecharging coil 24 is fully aligned with the receiving coil 109 at whichpoint the vibration becomes constant. Once the alignment has beenaccomplished and after a certain time period has elapsed, e.g., 30seconds, the processor 12 controls the vibrator 14 to stop the constantvibration.

At the same time, the processor 12 controls the speaker 16 to generate atone (e.g., beeps) having the same frequency and duration as thevibrator 14. In other words, the processor 12 can control the speaker 16to make a tone at the initial interval of 1-2 Hertz, generating thesound for 0.1 second each time. As the charger 210 comes closer to theIPG 102, the tone can correspondingly increase to 8-9 Hertz with thesame 0.1 second sound duration. When the charger 210 is fully aligned,i.e., the output of the sensor 38 has reached the threshold value, thenthe processor 12 controls the speaker 16 to generate a continuous toneto indicate that the charger 210 is now fully aligned with the IPG 102.Once the alignment has been accomplished and after a certain time periodhas elapsed, e.g., 30 seconds, the processor 12 controls the speaker 16to stop the continuous tone.

As can be appreciated, the vibrator 14 providing tactile feedback to thepatient can be very important because in certain environments, thepatient may not be able to hear the audible feedback from the speaker16.

In the IPG 102, the receiving coil 109 is inductively coupled to thecharging coil 24 when they are positioned near each other. The magneticfield from the charging coil 24 induces voltage in the receiving coil109. As an example, the induced voltage is an oscillating voltage with aswing of +3V to −3V for a peak to peak voltage of 6V. A resonance tank28 which includes a capacitor connected in series with the receivingcoil 109 comprises a resonance circuit whose resonance frequency istuned to the frequency of the magnetic field emanating from the chargingcoil 24. The induced voltage is rectified by a rectifier 30 to convertan oscillating voltage into a DC voltage. In the embodiment shown, therectifier 30 is a full wave voltage doubler rectifier so as to generatea 6V DC at its output. The IPG processor 104 such as a microcontrollercontrols a switch 32 connected across the rectifier 30 in parallel. Oneend of the switch 32 is connected to a power management circuit 34 whilethe other end is connected to ground Vss. In one embodiment, the switch32 is a MOSFET transistor that can be turned on or off by the processor104. Normally, the switch 32 is turned off including the time when thebattery 108 is being charged.

The power management circuit 34 receives the rectified DC voltage fromthe rectifier 30 and charges the rechargeable battery 108. Othercircuits 35 control the actual generation and controlling of the spinalcord stimulation signals.

Once the charger 210 and the implanted IPG 102 are aligned, the chargeris strapped to the body of the patient so that it is fixed relative tothe IPG and the charger starts charging the IPG battery 108. However,when the charger 210 continues to charge the battery 108 in the IPG 102even when it has fully charged, the extra induced power can potentiallydamage the various circuits in the IPG. To prevent such damage, thecharger 210 would need to turn off the power amplifier 22. Since thereis no active communication from the IPG 102 to the charger 210, it is achallenge to detect when the battery 108 of the IPG 102 has fullycharged.

According to another aspect of the present invention, a novel way ofdetecting the end-of-charge is disclosed. When the power managementcircuit 34 determines that the IPG battery 108 has been fully charged,it sends an end-of-charge signal to the IPG processor 104. A smallend-of-charge software is stored in an internal memory 105 of theprocessor 104 and is executed by the processor upon receiving theend-of-charge signal from the power management circuit 34. Under thecontrol of the stored end-of-charge software, the processor 104 turns onand off the switch 32 to electrically short the receiving coil 109 toground in a selected pattern. For example, the switch 32 could be turnedon and off at 1 Hertz for at least 3-5 times with a 50% duty cycle. Inother words, the switch 32 could be on for 0.5 second and off for 0.5second, and the on-off operation of the switch could be repeated atleast 3 times, preferably at least 5 times and most preferably at least10 times.

At the charger 210, a small end-of-charge detection software is storedin the internal memory of the processor 12 and is executed by theprocessor. The end-of-charge detection software continuously monitorsthe reflected impedance values from the sensor 38 for purposes ofdetecting an end-of-charge signal from the IPG 102. The processor 12 andthe internally stored end-of-charge detection software comprise anend-of-charge detection circuit. When the switch 32 from the IPG 102turns on and off repeatedly, the electrical short created by the switchcauses the output of the sensor 38 to go up and down in a predeterminedpattern according to the on-off switching pattern of the IPG switch. Inone embodiment, the predetermined pattern is a sine wave shape. Theprocessor 12 could detect the end-of-charge by recognizing that pattern.For example, if the on-off state of the switch 32 is repeated 10 times,then the processor 12 could count the number of times the sensor 38output rises above a threshold value. If the number is 8 or greater,then the processor could determine that the end-of-charge status of thebattery 108 has been reached.

Alternatively, the processor 104 could vary the current being receivedby the receiving coil 109 in a selected pattern which corresponds to thepredetermined pattern of the reflected impedance sensed by the sensor38. This could be accomplished, for example, by varying the amount of onor off state of the switch 32.

Once the processor 12 determines that the end-of-charge status has beenreached, it controls the vibrator 14 and the speaker 16 to outputtactile and audible signals which is indicative of the end-of-chargestatus of the IPG battery 108. For example, the vibration and beep couldlast for 0.1 second, two times a second for about 10 seconds.Thereafter, the processor 12 could turn off the current to the chargingcoil 24 and possibly turn itself off completely.

The charging frequency of the power amplifier 22 is at a set frequencywhich ensures maximum power transfer to the IPG 102. In the embodimentshown, the charging frequency is set at about 85 kHz which is theoptimum resonant frequency of the charger 210 and IPG 102 at an idealseparation distance of 15 mm.

However, the optimum operating frequency may change from the 85 kHz setfrequency depending on many factors such as the presence of metallicobjects near the charger 210, proximity and size of such objects and thelike. Presence of metallic objects can affect the optimum operatingfrequency by as much as a few kHz in either direction. Consequently,operating the charger at only one set frequency may limit the maximumachievable power transfer to the IPG 102.

According to another aspect of the present invention, a novel way ofoptimizing the charging frequency is disclosed. An optimization circuitselects an optimum frequency of a charging signal supplied to thecharging coil 24 based on evaluation of the reflected impedances of agroup of charging frequencies in a selected frequency range. Anoptimization software is stored in the internal memory 13 of theprocessor 12 and is programmed to be executed by the processor at settime intervals. For example, the optimization software is programmed tobe executed about every half a minute to about 5 minutes.

The optimization software, processor 12 and reflected impedance sensor38 comprise the optimization circuit. In short, the optimization circuitsweeps the charging frequencies within a small band of frequencies witha selected step size. At every frequency step, the charger 210 estimatesthe power transfer and moves to the next frequency step. At the end ofsweeping the frequency band, the charger will be set to the frequency atwhich the estimated power transfer is the highest. The sweeping of thefrequencies to find the optimum operating frequency by the optimizationsoftware will be performed once in about 30 seconds to about 5 minutesto ensure optimum operation. At each frequency step, the power transferis estimated based on the peak to peak voltage in the magnetic coil 24as measured by the reflected impedance sensor 38.

FIG. 24 is a detailed flowchart of the steps to optimizing the chargingfrequency by the optimization software. In step 54, the software sets asweep frequency range and initializes various variables. The selectedsweep frequency range in one embodiment is 80 kHz to 90 kHz. In step 54,an initial frequency is set to the lowest frequency in the range, andthe selected step interval is set at 100 Hz. Thus, sweeping across the80-90 kHz range takes 100 iterations.

In step 56, the reflected impedance sensor 38 continuously detects areflected impedance of the charging coil 24 and the optimization circuitreceives the detected reflected impedance values from the sensor. Theprocessor 12 under the control of the optimization software receives atleast several values from the sensor 38 and averages them. In oneembodiment, the processor stores 10 values from the sensor 38sequentially and then averages them to produce a current averagereflected impedance value.

In step 58, the processor 12 determines whether the current averagevalue is lower than the interim stored value which represents the lowestreflected impedance value during the sweep. If so, that means that atthe current frequency being evaluated, less current is being detected bythe sensor 38 as more current/power is being transferred to the IPG 102.Control then transfers to step 60 where the current frequency is set tothe interim optimal frequency. On the other hand, if the processor 12determines that the current average value is higher than the interimstored value, then at the frequency being evaluated, less current/poweris being transferred to the IPG 102 and control passes to step 62.

At step 62, the processor 12 determines whether there is any morefrequency to evaluate. If so, control passes to step 68. At step 68, thecurrent frequency is incremented by the selected step interval (e.g.,100 Hz) and the evaluation process of steps 54-68 are repeated for thenext frequency.

However, if the processor 12 determines that there are no morefrequencies to evaluate, control passes to step 64 where the optimumfrequency of the charging signal to the charging coil 24 is set to theinterim optimal frequency which corresponds to the lowest interim storedvalue. In step 66, once the optimal frequency has been set, theoptimization circuit waits for ‘m’ minutes and the entire evaluationprocess repeats starting from step 54. In one embodiment, “m” is between0.5 minute and 5 minutes. For example, “m” could be set to 2 minutes.

The entire frequency sweep from 80 kHz to 90 kHz can be done in lessthan 15 seconds, and preferably in less than 10 seconds.

The foregoing specific embodiments represent just some of the ways ofpracticing the present invention. Many other embodiments are possiblewithin the spirit of the invention. Accordingly, the scope of theinvention is not limited to the foregoing specification, but instead isgiven by the appended claims along with their full range of equivalents.

What is claimed is:
 1. A method for a wireless charger to automaticallytune an optimum frequency to inductively charge a rechargeable batteryof an implantable pulse generator (IPG) that generates spinal cordstimulation signals for a human body, the method comprising: applying aplurality of charging frequencies in a selected frequency range to acharging coil; detecting a reflected impedance of the charging coil foreach applied charging frequency; and selecting an optimum frequency ofthe charging coil.
 2. The method of claim 1, wherein selecting theoptimum frequency of the charging coil is based on the detectedreflected impedances of the plurality of charging frequencies
 3. Themethod of claim 2, further comprising receiving the detected reflectedimpedances of the plurality of charging frequencies by amicrocontroller, wherein the microcontroller selects the optimumfrequency based on the received impedances.
 4. The method of claim 1,further comprising periodically repeating the applying, detecting andselecting steps at a selected time interval.
 5. The method of claim 1,wherein the step of applying includes sweeping the plurality of chargingfrequencies in the selected frequency range from one end to the other ata selected interval to obtain the detected reflected impedances.
 6. Themethod of claim 1, wherein the step of selecting includes selecting, asthe optimum frequency, the charging frequency that results in thedetected reflected impedance representing the highest peak-to-peakvoltage of the charging coil.
 7. The method of claim 1, wherein the stepof detecting includes detecting the reflected impedance values from acurrent sensor coupled in series with the charging coil.
 8. The methodof claim 1, wherein the step of detecting includes detecting thereflected impedance values from a transformer having a primary windingin series with the charging coil and a secondary coil coupled to theprimary coil.
 9. The method of claim 1, wherein step of detectingincludes detecting the reflected impedance values from a rectifier thatrectifies the output of a current sensor coupled in series with thecharging coil.
 10. The method of claim 1, further comprising limitingthe current supplied to a class-E amplifier below a threshold value, thecurrent limiter connected between a power supply and the class-Eamplifier applying the plurality of charging frequencies to the chargingcoil.
 11. A method for a wireless charger to automatically tune anoptimum frequency to inductively charge a rechargeable battery of animplantable pulse generator (IPG) that generates spinal cord stimulationsignals for a human body, the method comprising: applying a plurality ofcharging frequencies in a selected frequency range to a charging coil;detecting a reflected impedance of the charging coil for each appliedcharging frequency; selecting an optimum frequency of the charging coilvia an optimization circuit, wherein the optimization circuit comprisesa optimization software, a processor, and a reflected impedance sensor.12. The method of claim 11, wherein selecting the optimum frequency ofthe charging coil is based on the detected reflected impedances of theplurality of charging frequencies
 13. The method of claim 12, furthercomprising receiving the detected reflected impedances of the pluralityof charging frequencies by a microcontroller, wherein themicrocontroller selects the optimum frequency based on the receivedimpedances.
 14. The method of claim 11, further comprising periodicallyrepeating the applying, detecting and selecting steps at a selected timeinterval.
 15. The method of claim 11, wherein the step of applyingincludes sweeping the plurality of charging frequencies in the selectedfrequency range from one end to the other at a selected interval toobtain the detected reflected impedances.
 16. The method of claim 11,wherein the step of selecting includes selecting, as the optimumfrequency, the charging frequency that results in the detected reflectedimpedance representing the highest peak-to-peak voltage of the chargingcoil.
 17. The method of claim 11, wherein the step of detecting includesdetecting the reflected impedance values from a current sensor coupledin series with the charging coil.
 18. The method of claim 11, whereinthe step of detecting includes detecting the reflected impedance valuesfrom a transformer having a primary winding in series with the chargingcoil and a secondary coil coupled to the primary coil.
 19. The method ofclaim 11, wherein step of detecting includes detecting the reflectedimpedance values from a rectifier that rectifies the output of a currentsensor coupled in series with the charging coil.
 20. The method of claim11, further comprising limiting the current supplied to a class-Eamplifier below a threshold value, the current limiter connected betweena power supply and the class-E amplifier applying the plurality ofcharging frequencies to the charging coil.